(A) Field of the Invention
This invention relates in general to a device suitable for heat and mass transfer, and, in particular, to a heat and mass transfer device usuable for cooling and oxygenating blood.
(B) Description of the Prior Art
In normal human circulatory systems, the venous blood enters the right heart cavities and is pumped to the lungs. There the blood is oxygenated after which it is returned to the left heart cavities. Arterial blood is then distributed to the body tissues. While passing through the capillary beds of the body, the arterial blood gives-up oxygen, picks-up carbon dioxide, and thus, becomes again the venous blood.
In an artificial cardiorespiratory device, e.g., a total heart-lung bypass, the entire systemic venous blood is prevented from entering the right heart cavities. Instead, it is drained by a mechanical pump into an artificial circuit outside the body and is oxygenated in an artificial gaseous exchange device. The "arterialized" blood is returned by another pump to the systemic arterial system through a cannula in a branch of the aorta. It then perfuses the various capillary beds, but is prevented from entering the blood-heart cavities by the closed aortic valve.
The concept of extracorporeal circulation as an aid to cardiac surgery originated as early as 1937. It was felt that a machine capable of performing the function of the heart and lungs would enable a surgeon to operate upon intra-cardiac abnormalities under direct vision in a relatively dry bloodless field. Meanwhile, the brain, the myocardium, the liver, the kidneys and other tissues would receive adequate flows of oxygenated blood in the artificial heart-lung machine.
The task of replacing adequately the heart and lungs of an average human adult by artificial devices, however, presented a formidable challenge. It took these early investigators more than 15 years of research and development before the gas exchange capacity which permitted extracorporeal circulation in man was achieved. Since about 1950, mechanical oxygenators have evolved from the stage of crude laboratory devices into rather dependable pieces of surgical equipment. There are now three basic classes of artificial lungs being used; the bubble oxygenators, the flm oxygenators, and the membrane oxygenators.
From the wealth of information gained in empirical development and laboratory trials, a concept of the "ideal oxygenator" has immerged. Even now, however, available heart-lung machines still fall short of this idea in many respects.
The qualities of an ideal oxygenator are usually described in terms of the efficient performance of a human lung as the gas exchanger and of the gentleness of the normal circulatory system in handling blood. Most persons skilled in the art agree on the following requirements:
1. The artificial lung must be able to oxygenate up to five liters of venous blood per minute to the range of 95-100 percent saturation.
2. Simultaneously, the gas exchange or mass transfer device must remove carbon dioxide in appropriate amounts so as to avoid either CO.sub.2 retention (respiratory acidosis) or CO.sub.2 depletion (respiratory alkalosis). A suitably large gas exchange capacity must be provided while keeping the blood content of the artificial lung within reasonable limits.
3. The mechanical process of gas exchange must be gentle enough to avoid destruction of formed elements of the blood or denaturation of plasma proteins.
4. The artificial lung must be of simple design, have as few components as practical and be of dependable construction so as to permit phase oxygenation over prolonged periods, easy cleaning and assembly, and reliable sterilization.
The lungs of an average human adult are able to introduce, according to the metabolic needs of the organisms, from 250 to 5,000 ml. of oxygen into the blood, and to remove about the same amounts of carbon dioxide. This performance requires a gas exchange surface of 80 to 100 m.sup.2 and a relatively steady blood volume of 700 to 900 ml. in the lungs, which is continuously removed and replenished at a flow rate varying from 4 to 30 liters per minute.
An artificial lung need only to equal the minimal performance of the human lung, because it is used for perfusion of the resting organism. It should be able, however, to bind up to 300 ml. of oxygen per minute to the venous blood.
Most artificial oxygenators, however, are limited to an oxygen uptake of between 150 and 250 ml./min. Unlike the natural lung, they require an increasingly large surface for gas exchange as the demand for oxygenated blood is augmented. Increased oxygen binding capacity is obtained only at the price of a larger blood volume or content in the oxygenator.
The priming volume of pump oxygenators sufficient in capacity to carry on the perfusion of human adults varies from 1200 to 6000 ml. The largest volumes of blood are almost prohibitive in terms of cost and time required for preparation. On the other hand, there is probably little practical advantage in reducing the priming volume below 1,000 ml. per human perfusion. Most surgeons find it comforting to have a 20 to 60 second supply of blood in the extracorporeal circuit, in case of massive bleeding or any other accident leading to temporary absence of venous return from the organism into the heart-lung machine. Seen in this light, satifying the need for increasingly large priming volumes, as higher flows are required, tends to keep constant the ratio of oxygenator blood content to blood flow. This is a safety factor which cannot be neglected under practical conditions.
During the natural process of blood oxygenation, as occurs in lungs or gills, there is no direct contact between blood and the ambient gas. A semipermeable membrane separates the blood from the oxygen present in the alveolar gas or in the surrounding water. The gas transfer is then evoked by process of molecular diffusion. Since many of the early difficulties encountered in using artificial lungs were associated with froth formation or fibran deposits at the "raw" blood-gas interface, artificial membranes were proposed as a means of protecting the blood from direct exposure to the atmosphere. The concept that respiratory gas transfer is compatible with physical separation of blood and oxygen has evolved into a number of different designs for artificial lungs during the last 10 years or so. In general, howeverever, a membrane oxygenator comprises essentially two membranes between which venous blood is conducted through a gas chamber. Oxygen and carbon dioxide are exchanged across the membranes and arterialized blood is led off.
In membrane oxygenators, a blood film and gaseous oxygen are separated by a semipermeable membrane. Gas transfer across the membrane depends upon the nature of membrane material, its thickness, surface and degree of hydration, and also upon the partial pressure difference of the diffusing gases on opposite sides of the membrane. Theoretical and experimental studies have indicated that CO.sub.2 transfer is primarily limited by the membrane barrier, while oxygen transfer is controlled by the thickness of and fluid flow character of the blood film, or in other words by the characteristics of the blood distributing system. Thus, the basic problem in the design of membrane lungs are logistic in nature. Some are associated with the membrane itself, others with the manifolding and distributing system for the blood.
When artificial membranes were first proposed, it was not immediately recognized that carbon dioxide transfer might be more difficult to carry out than oxygen transfer under the particular pressure conditions which prevail in an artificial lung. In the case of an artificial lung, since pure oxygen is used in the gas phase and since the partial pressure of CO.sub.2 in blood should not exceed 50 millimeters Hg., a ratio of at least 12:1 exists in favor of oxygen transfer in terms of pressure gradient. Accordingly, to counteract the partial pressure ratio and insure an equal transmission of CO.sub.2 and oxygen, the membrane should be 12 times more permeable to CO.sub.2 than to oxygen. Only then can a gas exchange ratio of 1 be maintained when the artificial lung is ventilated with pure oxygen.
Unfortunately, most synthetic membranes are only 4 to 5 times more permeable to CO.sub.2 than to oxygen, and this is insufficient for the transfer of equal volumes of oxygen and carbon dioxide. Thus, membrane lungs have to be designed in terms of CO.sub.2 release, and will therefore feature a considerable reserve in oxygen transfer capacity. In other words, the rate of permeation of CO.sub.2 is a bottleneck and dictates the area of the membrane required.
In the liquid and in the gaseous phase on either side of the membrane, CO.sub.2 diffuses very rapidly, so that these steps in the gas transfer hardly influence the overall rate of exchange. On the contrary, for the oxygen, factors such as blood film thickness or transit time are critical. It has been established experimentally that the oxygenating capacity of a membrane lung depends much more upon the blood distribution pattern than upon the membrane actually employed. This view is supported by a theoretical analysis of gas diffusion in a membrane lung. When highly permeable membranes are used in a blood oxygenator, the relative permeabilities of the blood film and of the membrane are such that the resistance of the blood film is the controlling factor. As in the case of film oxygenators which do not feature turbulent flow or continuous refilming, the length of the oxygen diffusion path is the rate limiting factor as soon as the thickness of the film exceeds a certain degree as compared to the size of an erythrocyte.
One can thus summarize the logistic problems of the membrane lung by stating that CO.sub.2 elimination from the blood depends upon the membrane area available, whereas oxygen uptake into the blood depends upon the design of the blood distributing system.
While logistics is a problem in the design of membrane oxygenators, another common problem associated therewith involves purging of trapped air initially in the blood film flow volume. The necessity for doing this is, of course, obvious as it would be extremely dangerous to allow trapped air to enter the blood stream. Various means have been devised to accomplish the desired purging; however, the primary ones involve tilting, tapping, and shaking of the oxygenator.
Hypothermia as an aid to cardiac surgery developed directly from the work of Bigelow et al. in 1950. Bigelow's fundamental observation was that the oxygen consumption of the warmblooded animal can be reversibly lowered by hypothermia without an oxygen debt being incurred. This reduction in tissue oxygen demand brought about by cooling prompted the idea that low flow perfusion might be usefully associated with hypothermia. The fall in blood temperature observed in most perfusion procedures is no longer feared, but rather is deliberately facilitated by heat-exchange devices. Light (33.degree.-35.degree. C.), moderate (26.degree.-32.degree. C.), and deep (0.degree.-20.degree. C.) hypothermia are commonly used in clinical practice.
In the early stages of clinical heart-lung machine development, temperature regulating devices were incorporated into the extracorporeal circuit to maintain the blood temperature within the normal physiological range. More recently, the success of hypothermic perfusions has led to the development of special heat exchangers of large calorie capacity, which are capable of modifying the temperature of the blood in both directions at a very rapid rate. Heating systems are sometimes required in perfusion equipment because of large metal, glass or plastic parts of heart-lung machines which present tremendous radiating surfaces and are responsible for most of the heat loss from the circulating blood.
Since there is considerable heat loss in all extracorporeal circuits various devices have been introduced for the purpose of maintaining the temperature of the perfused organism within the normal physiological range. Water baths, warm air, infrared radiant energy and heating wires have been used with varying degrees of success. Because of possible thermo injury to the red blood cells, the local temperatures of surfaces in contact with blood should never be permitted to exceed 45.degree. C.
Heretofore, others have disclosed devices comprising membranes having hemispherical protrusions or projections from the membrane for use in blood oxygenation. Such a membrane is shown in U.S. Pat. No. 3,684,097 to Mathewson et al. As disclosed in this patent, two membranes having hemispherical projections therefrom are used in combination, the membranes being disposed so that the projections of one membrane are interdigitated with those of a confronting membrane. This prevents membrane cohesion during storage.
Spiral wound membrane structures are also known, these being shown in U.S. Pat. Nos. 3,489,647 and 3,738,813.